In the past, the concept of volume responsiveness (sometimes also called “fluid responsiveness”) has been increasingly used in order to optimize the fluid management of patients, in particular of critically ill and/or anesthetised patients. According to this concept, a patient whose stroke volume, i.e. the amount of blood expelled by a ventricle of the heart during systole, increases significantly after fluid administration has “high” volume responsiveness. In contrast, a patient whose stroke volume hardly increases after fluid administration exhibits “low” or no volume responsiveness. Whether an increase in cardiac output will occur upon fluid administration depends mostly on the position of an individual patient on the so-called “Frank-Starling curve”. As schematically illustrated in FIG. 2, the Frank-Starling curve plots the preload against stroke volume, wherein by “preload” the volumetric pressure is meant that stretches the right or left ventricle of the heart to its largest geometric dimensions at the end of the diastole. Notably, the Frank-Starling curve is not linear but exhibits a characteristic concave shape, with the initial part being relatively steep and becoming progressively flat thereby reaching a plateau.
Achieving an increase in cardiac output usually results in an optimization of tissue perfusion which is a major goal of fluid therapy. As described for example by G. Gouvêa et al. in the British Journal of Anaesthesia (2009), 103: 238-43, “Evaluation of the pulse pressure variation index as a predictor of fluid responsiveness during orthotopic liver transplantation” and also by M. Cannesson et al. in the Journal of Clinical Monitoring and Computing (2011), 25: 45-56, “Pulse pressure variation: where are we today?”, if the patient is on the steep part of the Frank-Starling curve, cardiac output can be efficiently increased by administering fluid. However, if the patient is on the flat part of the Frank-Starling curve, no significant increase in cardiac output will be achieved by administering fluid. Fluid loading in this setting may even be hazardous for the patient and cause peripheral or pulmonary edema. Therefore, it is of major importance to reliably determine the position of the patient on the Frank-Starling curve, i.e. to determine the patient's volume responsiveness, before introducing fluid into the patient's circulation.
Since there is no straight-forward method available to measure the preload and the stroke volume of a patient's heart directly, i.e. to measure the patient's volume responsiveness, intensive studies have been carried out during the last decades to identify another indicator based on measurable parameters to predict the volume responsiveness of a patient.
As described for example by F. Michard et al. in Crit Care (2000), 4: 282-289, “Using heart-lung interactions to assess fluid responsiveness during mechanical ventilation”, in mechanically ventilated patients, the magnitude of the respiratory changes in left ventricle stroke volume can be used to assess volume responsiveness. Intermittent positive-pressure ventilation induces cyclic changes in the loading conditions of right and left ventricles. Mechanical insufflation decreases preload and increases afterload of the right ventricle. Right ventricular preload reduction is due to a decrease in the venous return pressure gradient that is related to the inspiratory increase in pleural pressure during mechanical positive pressure ventilation. The concomitant increase in right ventricular afterload is related to the inspiratory increase in trans-pulmonary pressure caused by the ventilation induced increase in airway pressure. Reduction in right ventricular preload and increase in right ventricular afterload both lead to a decrease in right ventricular stroke volume, which arises at its minimum at the end of the inspiratory period. The inspiratory impairment in venous return is assumed to be the major mechanism behind the inspiratory reduction in right ventricular stroke volume. The inspiratory reduction in right ventricular stroke volume subsequently leads to a further downstream decrease in left ventricular filling after a phase lag of two to three heart beats because of the blood's trans-pulmonary transit time. Thus, left ventricular preload reduction may induce a decrease in left ventricular stroke volume, which reaches its minimum during the mechanical expiratory period.
Interestingly, the cyclic changes in right ventricle preload induced by mechanical ventilation should result in greater cyclic changes in right ventricular stroke volume when the right ventricle operates on the steep rather than on the flat portion of the Frank-Starling curve. The cyclic changes in right ventricular stroke volume, and hence in left ventricular preload, should also result in larger cyclic changes in left ventricular stroke volume when the left ventricle operates on the ascending, steep portion of the Frank-Starling curve. Thus, the magnitude of the respiratory changes in left ventricular stroke volume which is a major determinant of systolic arterial pressure should be an indicator of volume responsiveness. Therefore, it has been proposed to analyse the respiratory changes in systolic pressure by calculating the difference between maximum and the minimum value of systolic pressure over one single respiratory cycle of a mechanically ventilated patient. This difference has been called “systolic pressure variation” (SPV).
Furthermore, it has been recently proposed that cardiac volume responsiveness may be assessed in a more sophisticated way by calculating the arterial “pulse pressure variation” (PPV). “Pulse pressure” (PP) has been defined as the difference between systolic and diastolic pressure within one single respiratory cycle. Notably, the pulse pressure is almost directly proportional to the stroke volume of the left ventricle. Conventionally pulse pressure variation is calculated using the following formula:
  PPV  =                              PP                      ma            ⁢                                                  ⁢            x                          -                  PP                      m            ⁢                                                  ⁢            i            ⁢                                                  ⁢            n                                                1          2                ⁢                  (                                    PP                              m                ⁢                                                                  ⁢                ax                                      +                          PP                              m                ⁢                                                                  ⁢                i                ⁢                                                                  ⁢                n                                              )                      *    100    ⁢    %  wherein PPmax and PPmin are the maximal and minimal pulse pressure, respectively, within one single respiratory cycle. Notably, during one single respiratory cycle the heart usually beats several times. For example, if the heart beats six times during one respiratory cycle, PPmax and PPmin for six heart beats may be measured. To calculate the pulse pressure variation, the single maximum pulse pressure value and the single minimum pulse pressure value for the entire sequence of six subsequent heart beats have to be determined.
Adversely, both previously described approaches for determining an indicator representative for volume responsiveness, i.e. the SPV-approach and PPV-approach, strictly require to reliably measure blood pressure values associated with each individual heart beat in the course of the detection period, e.g. during one respiratory cycle comprising approximately six heart beats. However, due to frequently occurring artefacts or arrhythmias of the heart, not all of the measured values actually reflect the patient's heart-lung interaction in terms of volume responsiveness. This phenomenon makes both above approaches prone to errors. In fact, analysis of the respiratory changes in arterial pressure is hardly possible in patients with cardiac arrhythmias.
A well-known and convenient way to determine the arterial blood pressure of a patient is to use the so-called “oscillometric non-invasive blood pressure measurement method”. By that method, the pressure in a pressure cuff, which is usually applied to the patient's arm, is continuously increased or decreased. For example, the pressure in the pressure cuff may be initially set to a value well above the systolic pressure of the patient, and may then be continuously decreased to a value below the diastolic pressure of the patient. Hereby, the pressure in the pressure cuff is continuously decreased over a time period which corresponds to a plurality of heart beats. A manometer is connected to the pressure cuff which not only indicates the continuously decreasing pressure applied to the pressure cuff, but, in addition (due to the principle action=reaction) also indicates the pressure oscillation based on the varying pulse signal, i.e. pulse amplitude and waveform. Further, the corresponding result of a single heart beat in any artery sensed by any method will be called a “pulse”. By plotting exclusively such cyclic pressure variations, i.e. oscillations, indicated by the manometer over the time, the oscillation amplitude is not constant but rather bell shaped. The maximum oscillation is usually reached when the pressure applied to the pressure cuff is somewhere in the middle between the systolic and the diastolic pressure of the patient. In other words, the sensitivity of the manometer with respect to the pressure oscillations caused by the heart beats is at its maximum whenever the cuff pressure substantially corresponds to an intermediate value between the systolic and the diastolic pressure (or is slightly below that intermediate value).
A pressure cuff being connected to a manometer is schematically shown in FIG. 3. The pressure cuff is applied to a patient's arm and indirectly, through skin, fat, muscles and inter-connective tissue, exerts pressure on an artery. An electrocardiogram (ECG) signal over the time is schematically shown in FIG. 4a. The pressure in the pressure cuff that is detected as a function of time by the manometer is schematically shown in FIG. 4b. The plotted pressure signals in FIG. 4b represent superposition of the continuously decreasing pressure applied to the pressure cuff on one hand, and cyclic pressure oscillations caused by the pulses of the patient on the other hand. In this example, the pressure in the pressure cuff is continuously decreased from a value above the systolic pressure Psys of the patient to a value below the diastolic pressure Pdia of the patient, e.g. via a not shown valve. Notably, instead of continuously decreasing the pressure in the pressure cuff, it is also possible to continuously increase the pressure in the pressure cuff. By another plot, exclusively the pressure oscillations detected by the manometer are schematically illustrated in FIG. 4c. These pressure oscillations are illustrated as oscillating around an average value. As can be seen from FIG. 4c, the amplitude of the pressure oscillations is not constant but has a maximum when the pressure in the pressure cuff substantially corresponds to the 50% intermediate value between the systolic blood pressure and the diastolic blood pressure of the patient.
As noted previously, by the above described oscillometric non-invasive blood pressure measurement method, the cuff pressure is continuously increased or decreased. Accordingly, no constant measurement conditions can be applied and, thus, this method only allows for determining a single systolic pressure value and a single diastolic pressure value. A reliable determination of the individual systolic and diastolic pressure values associated with each single individual heart beat cannot be provided by that blood pressure measuring approach. Consequently, such an approach does not reliably allow for determining the volume responsiveness of the patient via the SPV-approach or the PPV-approach.
EP 0 078 090 A1 teaches a non-invasive blood pressure measurement method that allows for determining pulse pressure variation. According to this method, a fluid-filled pressure cuff is permanently attached to a patient's finger. A light source and a light detector are integrated in the pressure cuff, the light source and the light detector forming part of a photo-electric plethysmograph. The cuff pressure is controlled—via an electric pressure valve—by the plethysmographic signal in closed-loop operation, so that the arterial volume in the finger is maintained at a value to be pre-adjusted. Measuring the pressure in the pressure cuff, thus, allows for determining the arterial blood pressure of the patient.
However, permanently pressing the sensors to the patient's finger negatively affects blood circulation and can cause severe tissue damage or even necrosis of the finger. Furthermore, the finger is relatively remote from the patient's heart, and the diameter of arterial vessels in the finger is relatively small compared to the diameter of arterial vessels close to the heart. Due to interference effects caused by pressure reflections occurring when the diameter of arterial vessels abruptly changes, e.g. when arterial vessels branch, the blood pressure wave form measurable at the finger only imprecisely corresponds to the arterial blood pressure of interest, namely the central arterial blood pressure and waveform closest to the patient's heart.
It is therefore an object of the present disclosure to provide a simple and robust method or means for reliably determining an indicator representative for the patient's volume responsiveness. In particular, it is an object of the present disclosure to provide a method or means for determining volume responsiveness of a patient, which method or means minimizes the negative impact (exerted e.g. by artefacts or arrhythmias of the patient's heart) on the indicator to be determined. The method or means according to the present disclosure should preferably also reliably be implemented on the basis of oscillometric non-invasive blood pressure measurement methods known in the art.